Diabetes mellitus is a disease of major global importance, and its frequency of incidence has been increasing at almost epidemic rates. The worldwide prevalence in 2006 was 170 million people, and this number is predicted to at least double over the next 10-15 years. Diabetes is generally characterized by a chronically raised blood glucose concentration (hyperglycemia), due to a relative or absolute lack of the pancreatic hormone, insulin. In a normal (non-diabetic) subject, pancreatic islet cells (beta cells) continuously sense the blood glucose levels and consequently regulate insulin secretion to maintain near constant levels. However, diabetic patients lack this capability.
Much of the burden of the disease to the patient and to health care resources is due to long-term tissue complications, which can affect both the small blood vessels (microangiopathy, causing eye, kidney and nerve damage) and the large blood vessels (causing accelerated atherosclerosis, with increased rates of coronary heart disease, peripheral vascular disease and stroke). There is now evidence that morbidity and mortality of diabetic patients is related to the duration and severity of hyperglycemia. In theory, maintaining normal blood glucose levels by hormone replacement therapy using insulin injections and/or other treatments in diabetes might be able to prevent complications. However, near-normal blood glucose levels can be quite difficult to achieve and maintain in many patients, particularly in those having Type 1 diabetes. In these patients, blood glucose concentrations can vary fairly quickly between very high (hyperglycemia) and very low (hypoglycemia) levels in an unpredictable manner.
Many diabetic patients currently measure their own blood glucose several times during the day by using finger-prick capillary samples and applying the blood to a reagent strip for analysis in a portable glucose meter. The discomfort involved with these tests can often lead to poor patient compliance. Testing cannot be performed while sleeping and while the subject is occupied. In addition, the readings do not give information regarding the trends in glucose levels, but rather provide only discrete readings, taken at large time intervals between the measurements. Therefore continuous glucose monitoring would be advantageous, providing essentially continuous glucose readings by performing discrete measurements, at a very high frequency.
An electrochemical glucose sensor is described in U.S. Pat. No. 6,612,111 assigned to Lifescan Inc., which is hereby incorporated by reference herein. Today, the majority of available electrochemical glucose sensors are enzyme-based. The detection principle of these sensors is based on the monitoring of the enzyme-catalysed oxidation of glucose. These include glucose sensors use amperometric or potentiometric operating principles.
The enzymatic reaction that occurs in the majority of these sensors is catalyzed by glucose oxidase (GOX). During this reaction, oxygen and glucose yield gluconic acid and hydrogen peroxide. Glucose oxidase acts temporarily as an electron acceptor, where it is first reduced to an inactive state and subsequently is reactivated by the reduction of oxygen to hydrogen peroxide. The glucose concentration is transformed into a detectable signal, which is proportional to the glucose level and which is generally measured by amperometric methods.
An enzyme-coated working electrode can serve as the sensor transducer, which is where electrochemical oxidation or reduction takes place. A counter electrode can be paired with the working electrode. A current of opposite sign passes through the two electrodes. The intensity of the current is a function of the concentration of electro-active glucose. An increased surface area between the analyte sensing layer (containing the enzyme) and the working electrode can enable enzyme loading, which is necessary for overcoming degradation of the enzyme as the reaction proceeds. The increased surface area can also enable enhanced electron transfer between the enzyme active site and the sensor transducer, thus improving the sensor performance.
Several ambulatory insulin infusion devices are currently available on the market. The first generation disposable devices configured as syringe-type reservoir are described in 1972, by Hobbs, in U.S. Pat. No. 2,631,847, and in 1973, by Kaminski, in U.S. Pat. No. 3,771,694, and later by Julius, in U.S. Pat. No. 4,657,486, and by Skakoon, U.S. Pat. No. 4,544,369, each of which is hereby incorporated by reference herein. These devices are generally quite large and heavy due to their spatial design and the relatively large driving mechanism of the syringe and the piston. This relatively bulky device has to be carried in a patient's pocket or attached to the belt. Consequently, the fluid delivery tube can be quite long, in some cases grater than 40 cm, to permit needle insertion in remote sites of the body. Such uncomfortable, bulky fluid delivery devices can be rejected by many diabetic insulin users, because of their negative impact on the performance of regular activities, such as for example sleeping and swimming. Furthermore, some more self-conscious users, such as for example teenagers, are likely to reject the use of such a device because of the potential negative body image that might result from using it. In addition, the long delivery tube can exclude some potential remote insertion sites, such as for example the buttocks and the extremities.
To avoid potential disadvantages associated with tubing, a second generation of pumps based on a new concept has been devised. These pumps can include a housing having a bottom surface adapted for attaching to the user's skin, a reservoir disposed within the housing, and an injection needle in fluid communication with the reservoir. These skin adherable devices are generally disposed of every 2-3 days similarly to the infusion sets employed in the pumps of the first generation. Devices of this type have been described by Schneider, in U.S. Pat. No. 4,498,843, Burton in U.S. Pat. No. 5,957,895, Connelly, in U.S. Pat. No. 6,589,229, and by Flaherty in U.S. Pat. No. 6,740,059, each of which is incorporated by reference herein. Other configurations of skin adherable pumps are disclosed in U.S. Pat. Nos. 6,723,072 and 6,485,461, which are also incorporated by reference herein. The pumps described in these references are generally designed as a single piece and remain adhered to the user's skin for the entire usage duration. The needle emerges from the bottom surface of the device and is fixed to the device housing.
Another fluid delivery device is described in international patent application no. PCT/IL06/001276, which is currently and was at the time of the development of the current subject matter commonly owned with the present application and is incorporated by reference herein. This device is configured as a miniature portable, programmable, skin-adherable fluid dispenser, which does not employ long tubing. The device includes two parts: a disposable part and a reusable part. The reusable part includes the necessary electronic components along with driving and pumping mechanisms. The disposable part includes reservoir for therapeutic fluid, short delivery tube and exit port. This fluid delivery device can also include a remote control unit that allows data acquisition, programming, and user inputs. Even after connection of the reusable and disposable parts, the assembled device has a very thin dimension, rendering the whole device inexpensive, light, and discrete.